The present invention relates to analytic chemical measurements, and more particularly to biosensor surfaces and related methods.
Analytical instruments such as biosensors are well established as a means of recording the progress of biomolecular interactions in real time. Biosensors are analytical instruments that employ a variety of transduction technologies in order to detect interactions between biomolecules. A particularly effective evanescent field based technology, known as surface plasmon resonance (SPR), exploits the behavior of light upon reflection from a gold-coated optical substrate.
SPR is an optical technique that enables real-time monitoring of changes in the refractive index of a thin film close to the sensing surface. The evanescent field decays exponentially from the gold surface and falls to one third of its maximum intensity at approximately 300 nm from the surface. Hence, the SPR technique is sensitive to surface refractive index changes and is almost completely insensitive to refractive index variations which occur more than 300 nm from the surface. An integrally-formed miniature SPR transducer has previously been described in U.S. Pat. No. 5,912,456. In this device a photodiode array (PDA) simply records the intensity of the reflected light, from an light emitting diode (LED), over a range of angles. Refractive index changes within the penetration depth of the evanescent field give rise to corresponding angular changes in the position of the SPR reflectance minimum. This change in resonance angle is followed by tracking the change in the PDA pixel position of the reflectance minimum. A minimum tracking algorithm is employed to continuously monitor the position of this minimum as it traverses the photodiode array and the pixel position is then related to a refractive index value. The current configuration of this device possesses three SPR active sensing regions per sensor enabling multichannel operation with real-time reference subtraction. Alternative configurations can allow as many as 100 or more SPR sensing regions.
The delivery of samples to the SPR active sensing regions is made possible by creating flow channels that cover the active sensing regions. Each flow channel possesses an inlet and outlet to allow for the flow of buffer, or samples, over the SPR active sensing regions.
In order to provide specificity to SPR, a biomolecule is typically immobilized near the gold surface. This immobilized biomolecule is referred to as a ligand. The immobilized ligand usually possesses binding specificity for another biomolecule contained in a sample, this other biomolecule is referred to as the analyte. The strength of this binding is given by the affinity constant (K) which is simply the ratio of the association rate constant (ka) divided by the dissociation rate constant (kd). It is possible to measure these constants because an SPR-based biosensor records the progress of binding and dissociation events in real time. To a large degree the performance of the biosensor is dictated by the properties of the surface preparation. Since the specificity of SPR-based biosensors is dictated by the choice of ligand to be immobilized at the sensing surface, it is possible to detect a wide range of analytes by simply choosing the appropriate ligand. However it is important that the ligand is immobilized in such a way as to retain its solution phase characteristics of solubility, accessibility, and analyte binding activity.
There are several properties that an attachment layer, with immobilized biomolecules, must possess in order to ensure optimal biosensor performance. These properties may vary depending on the requirement of a particular application. The following is a list summarizing the properties of an SPR biosensor surface that should be considered.
(1) Attachment Stability. The surface should possess ligands linked via stable bonds, such as covalent bonds, to ensure that leaching of the ligand from the surface during an application does not occur. It is important that the analyte binding capacity of the surface remains constant throughout the duration of a measurement or set of measurements.
(2) Three-Dimensional Accessibility. The ligand is best linked in such a way as to be fully accessible to the analyte. When the ligand is directly linked to a planar surface, for example, then three-dimensional access is not possible. Such a situation can alter the activity of the ligand and can thereby interfere with the extraction of accurate kinetic data.
(3) Capacity. Some biosensing applications require a large amount of ligand to be held by the surface. Examples of such applications include direct binding assays to determine the concentration of analyte down to very low levels. Some biosensing applications require low to medium levels of bound ligand to be held by the surface. Kinetics studies, equilibrium studies and competitive assays fall into this category because very high levels of ligand can impose mass transport limited conditions and other hindrances.
(4) Resistance to Non-Specific Binding (NSB). Binding of sample components other than analyte to the biosensing surface is referred to as non-specific binding. The performance of any biosensing technology can be reduced by the occurrence of NSB and so the NSB properties of the surface immobilization chemistry are important.
Hydrophobic sites at the surface commonly give rise to an increase in non-specific binding because physisorption of proteins to surfaces is known to be mediated by hydrophobic interactions. Also, an excess of charged groups also generally increases the probability of non-specific binding. For example, some proteins possess a net positive charge at neutral pH and will tend to associate with negatively charged surfaces.
(5) Uniform Distribution of Binding Sites. The surface should provide a uniform distribution of ligand across the sensing surface. But for highly surface-sensitive techniques such as SPR, it is also important to maintain a uniform distribution of immobilized ligands within the evanescent field penetration depth (i.e., above the surface). A non-uniform distribution in this dimension will also create problems for kinetic models since it is very difficult to account for such heterogeneity, and the analysis, therefore, becomes less robust. Specifically, typical kinetic binding models assume that analyte binds to all available binding sites homogenously and not as an inward moving analyte binding front.
(6) Monovalent Linkage. Multiple linkages between the ligand and the surface can cause steric hindrance by obstructing analyte binding sites. Thus, multivalent linkage can cause a fraction of immobilized ligand to be rendered less accessible than others. This will not only lower surface capacity, but more importantly will create a population of differentially accessible analyte binding sites. Meaningful kinetic data cannot be recorded under these conditions. Ligand binding sites must be fully accessible and this is best accomplished by ensuring a single linkage with the surface.
(7) Diffusion Limitations. Transport of analyte to the immobilized ligand at the sensing surface is a function of the diffusion rate of the analyte, the liquid flow rate, and the dimensions of the flow cell. High mass-transport rates typically yield more accurate kinetic data. However, mass transport of analyte may be distorted by the presence of three-dimensional structures, such as hydrogels, at the sensing surface. The analyte diffusion rate decreases within the hydrogel, thus compromising kinetic analyses.
(8) Chemical Resistance. The gold film should be insulated from the aqueous environment using a chemically-resistant thin film. Analytes are sometimes poorly soluble in aqueous buffers and the addition of solvents such as DMSO, or DMF, often enables solublization. Therefore that thin film must tolerate short-term exposure to such solutions. Resistance to common buffers, weak acids, weak bases, surfactants, and denaturants (e.g. 6M guandine-HCl, pH 2.0) is also desired.
SPR technologies have been available commercially for over a decade, and a considerable body of knowledge now exists as evidenced by thousands of peer-reviewed journal articles. A review of the literature indicates that kinetic analysis remains the mainstay of SPR applications, but other applications, such as concentration measurement, epitope mapping, interaction dynamics, ligand fishing, and affinity analysis are also common. Since the properties of the film that allows for ligand attachment at the biosensing surface is a critical component for all these applications, and because different applications have different surface demands, many approaches to surface attachment chemistries have been developed.
The most basic approach that has been applied to SPR involves directly attaching ligands to the gold surface and without the benefit of an intermediate film. While this approach is simple enough, its limitations are numerous. First of all, since this is a planar two-dimensional surface, the maximum amount of ligand that can be attached is a monolayer (approximately 3000 RU for an antibody, for example, where 1 RU is equivalent to a 1×10−6 change in refractive index or a surface coverage of approximately 1 pg/mm2). In the case where the ligand is a protein, which is most common, most of the protein attached in this manner will denature on the gold surface thereby rendering it inactive. And even those proteins that manage to remain active will suffer from steric hindrance because this binding is taking place directly on a planar surface. This will further reduce the ability of the analyte to access the active binding regions of the ligand. So, this technique produces limited loading and poor activity due to denaturation and steric hindrance.
U.S. Pat. No. 5,242,828 describes an improvement over direct ligand attachment to the gold by construction of a self-assembled monolayer (SAM) of alkane thiolates that chemisorb onto a gold surface to from a pseudo-crystalline monolayer as an intermediate layer for biosensing. Here the thiol end of the molecule binds to the gold and the film organizes itself into a self-assembled monolayer. This type of film provides an effective insulating layer that both protects the gold from harsh reagents and also protects the ligand from the denaturing effects of direct binding to a gold surface. In addition, the non-thiol end of the alkane thiol can be functionalized with hydrophilic groups to promote bio-compatibility and also with reactive groups to allow for covalent attachment. The practice of applying alkane thiols to gold surfaces uniformly can be tricky because any contamination, which forms readily on the gold upon exposure to air, will result in non-uniform SAM formation.
A modification of the alkane thoiolate SAM approach is described in U.S. Pat. No. 5,436,161. This describes the grafting of a porous matrix (e.g., hydrogel) onto the alkanethiolated surface, where that hydrogel contains both charged groups for preconcentrating oppositely charged ligand molecules into the gel and reactive groups that allow for covalent attachment of ligand molecules to the gel. This matrix coating is typically ˜100 nm thick, and due to its three dimensional nature, allows for up to ten times the binding capacity of a normal planar surface coating. The use of hydrogels on the sensing surface was first proposed by Liedberg et al. in 1983. The most popular hydrogel matrix used today is a carboxymethylated dextran-based hydrogel. Its high density of carboxyl groups throughout the three-dimensional gel is suspected of causing non-specific binding when used with some biological systems. In addition, the use of a high-density anchoring matrix in a biosensor also creates the potential for several well-documented matrix-related artifacts in kinetic studies as described below.
Random grafting of the dextran polymer, at any point along the chain, to the surface is unavoidable in the fabrication of hydrogel-based coatings. Non-uniform ligand binding with poor penetration of analyte to ligands located close to the surface may occur due to an increased density of polymer chains closer to the surface. This effect becomes more pronounced for high molecular weight analytes. Concerns related to the increasing density of the gel near the surface, have been noted. Indeed, D. Hall, Use of Optical Biosensors for the Study of Mechanistically Concerted Surface Adsorption Processes, 288 Anal. Biochem. 109-125 (2001), concluded that “As the existence of possible gel-induced partition and transport limitations cannot, in consideration, be divorced from the external transport processes as well as from consideration of the fundamental properties of the interacting system (intrinsic chemical reaction kinetics, ligate and ligand size, the concentration of ligate and immobilized receptor), the provision of clear guidelines for experimentation is difficult”. Also, the potential kinetic artifacts related to a variable density of ligand throughout the depth of the evanescent field are to be considered.
In a carboxymethylated dextran coating, for example, both the negative charge and reactive groups are distributed throughout the hydrogel since there is no way to specifically position them, and while this results in a highly effective preconcentration and immobilization of ligand, this process also allows the ligand to be bound by more than one linkage site and this is quite problematic. In fact it is likely that as many as three or four linkages are possible. Löf{dot over (a)}cs at al., Dextran Modified Gold Surfaces for Surface Plasmon Resonance Sensors: Immuno-reactivity of Immobilized Antibodies and Antibody-Surface Interaction Studies, 1 Colloids and Surfaces B, Biointerfaces, 83-89 (1993), found lower specific binding activities of ligands that were immobilized onto the carboxymethylated hydrogel via random amine coupling. This was attributed to multipoint attachment to the hydrogel. In cases where the occurrence of multipoint attachment was reduced (accomplished by using very high density ligand coupling), the analyte binding activity was found to be as high as 75%. But since low immobilization levels are required for most applications, this tendency to form multiple cross-linkages is quite problematic for hydrogel matrices due to their inherently high densities.
Furthermore, because there is charge throughout the hydrogel, and because the presence of charge affects the gel's structure, there is a tendency for the hydrogel to cause shifts in the baseline during biosensor measurements. These shifts result from the fact that changes in pH or ionic strength tend to change the density of charged hydrogels, and such density fluctuations give rise to undesirable shifts in the baseline response.
So, while the hydrogel matrix described in U.S. Pat. No. 5,436,161 does enable very high ligand loading, limitations of this method include analyte transport difficulties, multivalent attachment, ligand gradients and baseline instability.
Other work using carboxymethylated hydrogels as anchoring layers for biosensor surfaces includes Piehler et al., A High Density Poly(ethylene glycol) Polymer Brush for Immobilization on Glass Type Surfaces. 15 Biosensors & Bioelectronics, 473-481 (2000), which concluded, “These layers reduce non-specific binding, but proportionally interfere with the binding event and can also affect the sensitivity of the transducer”. Consequently, Piehler et al. replaced the dextran hydrogel with densely packed polyethylene glycol (PEG) chains grafted onto a planar surface. Similar to the results obtained using the technique of U.S. Pat. No. 5,242,828 non-specific binding was also found to be extremely low, but only a monolayer of ligand could be bound.
In addition, Piehler at al. had previously found that PEG chains did not effectively prevent non-specific binding when grafted at less than maximal densities. This is in agreement with the conclusions of Sheth et al., Measurements of Attractive Forces between Proteins and End-grafted Poly(ethylene glycol) Chains. 94 Proc. Natl. Acad. Sci., 8399-8402 (1997), working on attractive forces between PEG and proteins, which concluded “ . . . the activation energy for protein polymer adhesion increased with the polymer grafting density. Denser layers thereby increase not only the diffusional barrier but also the energy required to form attractive protein-polymer bonds”. Simply put, the conclusion was that the higher the density of the PEG layers, the more effective the film became at resisting non-specific binding.
Thus, current SPR sensors have problems related to their surface coatings. Planar two-dimensional surfaces allow only for monolayer ligand surface coverage. Matrix-based hydrogels allow for up to ten times higher ligand loading, but they also impart artifacts in kinetic data related to limited transport through the gel, ligand density variations, charge induced density changes, and multivalent attachment.